U.S. patent application number 13/028133 was filed with the patent office on 2011-08-18 for reducing data acquisition, power and processing for hemodynamic signal sampling.
Invention is credited to Robert G. Turcott.
Application Number | 20110201946 13/028133 |
Document ID | / |
Family ID | 43741723 |
Filed Date | 2011-08-18 |
United States Patent
Application |
20110201946 |
Kind Code |
A1 |
Turcott; Robert G. |
August 18, 2011 |
REDUCING DATA ACQUISITION, POWER AND PROCESSING FOR HEMODYNAMIC
SIGNAL SAMPLING
Abstract
Methods, systems and devices are provided for reducing the
amount of data, processing and/or power required to analyze
hemodynamic signals such as photoplethysmography (PPG) signals,
pressure signals, and impedance signals. In response to detecting a
specific event associated with a cyclical body function, a
hemodynamic signal is continuously sampled during a window
following the detecting of the specific event, wherein the window
is shorter than a cycle associated with the cyclical body function.
The hemodynamic signal is then analyzed based on the plurality of
samples. This description is not intended to be a complete
description of, or limit the scope of, the invention. Other
features, aspects, and objects of the invention can be obtained
from a review of the specification, the figures, and the
claims.
Inventors: |
Turcott; Robert G.;
(Mountain View, CA) |
Family ID: |
43741723 |
Appl. No.: |
13/028133 |
Filed: |
February 15, 2011 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10894962 |
Jul 19, 2004 |
7909768 |
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13028133 |
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Current U.S.
Class: |
600/484 ;
600/300; 600/508 |
Current CPC
Class: |
A61B 5/1455 20130101;
A61B 2560/0209 20130101 |
Class at
Publication: |
600/484 ;
600/300; 600/508 |
International
Class: |
A61B 5/0205 20060101
A61B005/0205; A61B 5/00 20060101 A61B005/00; A61B 5/024 20060101
A61B005/024 |
Claims
1. A method for reducing the amount of data required to analyze a
hemodynamic signal, comprising: (a) in response to detecting a
specific event associated with a cyclical body function,
continuously sampling the hemodynamic signal during a window
following the detecting of the specific event, to thereby produce a
plurality of samples, wherein the window is shorter than a cycle
associated with the cyclical body function; (b) repeating step (a)
a plurality of times such that a plurality of samples is produced
for each of a plurality of said windows; and (c) analyzing the
hemodynamic signal based on the plurality of samples, wherein said
analyzing step comprises averaging the plurality of samples
produced for each of the plurality of said windows to thereby
produce an average value for each of said windows.
2. The method of claim 1, wherein step (c) includes detecting a
peak-to-peak amplitude based on the plurality of samples.
3. The method of claim 1, wherein the continuous sampling of the
hemodynamic signal begins a fixed delay after a specific cardiac
event.
4. The method of claim 1, wherein the cyclical body function
comprises heart beat, and wherein the cycle associated with the
cyclical body function comprises a cardiac cycle.
5. The method of claim 4, wherein the continuous sampling is
performed at a rate of about 20 Hz or greater.
6. The method of claim 5, wherein the specific event comprises one
of the following: sensing a ventricular contraction; a ventricular
pace; sensing an atrial contraction; and an atrial pace.
7. The method of claim 1, wherein the cyclical body function
comprises respiration, and wherein the cycle associated with the
cyclical body function comprises a respiratory cycle.
8. The method of claim 1, wherein the continuous sampling is
performed at a rate of about 1 Hz or greater.
9. The method of claim 8, wherein the specific event comprises one
of the following: expiration; and inspiration.
10. The method of claim 1, wherein the hemodynamic signal comprises
a photoplethysmography (PPG) signal produced using a light source
and a light detector.
11. The method of claim 1, wherein the hemodynamic signal comprises
a pressure signal produced using a pressure transducer.
12. The method of claim 1, wherein the hemodynamic signal comprises
an arterial pressure signal produced using a pressure
transducer.
13. The method of claim 1, wherein the hemodynamic signal comprises
an atrial pressure signal produced using a pressure transducer.
14. The method of claim 1, wherein the hemodynamic signal comprises
a ventricular pressure signal produced using a pressure
transducer.
15. The method of claim 1, wherein the hemodynamic signal comprises
an electrical impedance signal produced using electrodes.
16. A system for reducing the amount of data required to analyze a
hemodynamic signal, comprising: one or more sensors configured to
produce one or more hemodynamic signals, wherein the hemodynamic
signals comprise at least a photoplethysmography (PPG) signal; a
light source and a light detector to produce the PPG signal; and a
computer readable medium comprising computer executable
instructions that when executed by a processor cause the processor
to: (a) in response to detecting a specific event associated with a
cyclical body function, continuously sample the one or more
hemodynamic signals during a window following the detecting of the
specific event, to thereby produce a plurality of samples, wherein
the window is shorter than a cycle associated with the cyclical
body function; (b) repeat step (a) a plurality of times such that a
plurality of samples is produced for each of a plurality of said
windows; and (c) analyze the one or more hemodynamic signals based
on the plurality of samples, wherein said analyzing step comprises
averaging the plurality of samples produced for each of the
plurality of said windows to thereby produce an average value for
each of said windows, wherein the cyclical body function comprises
respiration, and wherein the cycle associated with the cyclical
body function comprises a respiratory cycle.
17. The system of claim 16, wherein the continuous sampling is
performed at a rate of about 1 Hz or greater.
18. The system of claim 16, wherein the continuous sampling is
performed at a rate of about 1 Hz or greater and wherein the
specific event comprises one of the following: expiration; and
inspiration.
19. A system for reducing the amount of data required to analyze a
hemodynamic signal, comprising: a sampler configured to
continuously sample the hemodynamic signal during a window
following a detection of a specific event associated with a
cyclical body function, to thereby produce a plurality of samples,
wherein the window is shorter than a cycle associated with the
cyclical body function; a processor configured to analyze the
hemodynamic signal based on the plurality of samples wherein the
processor averages the plurality of samples produced for the window
to thereby produce an average value for the window, wherein the
hemodynamic signal comprises a photoplethysmography (PPG) signal;
and a light source and a light detector to produce the PPG signal,
wherein the cyclical body function comprises respiration, and
wherein the cycle associated with the cyclical body function
comprises a respiratory cycle.
20. The system of claim 19, wherein the continuous sampling is
performed at a rate of about 1 Hz or greater and wherein the
specific event comprises one of the following: expiration; and
inspiration.
Description
PRIORITY CLAIM
[0001] This application is a Divisional Application of and claims
priority and other benefits from U.S. patent application Ser. No.
10/894,962 (Attorney Docket No. A04P3010-US3), filed Jul. 19, 2004,
entitled "REDUCING DATA ACQUISITION, POWER AND PROCESSING FOR
HEMODYNAMIC SIGNAL SAMPLING", incorporated herein by reference in
its entirety.
FIELD OF THE INVENTION
[0002] Embodiments of the present invention relate to reducing the
amount of data, processing and/or power required to analyze
hemodynamic signals such as photoplethysmography signals and
arterial pressure signals.
RELATED APPLICATIONS
[0003] The present application is related to the following commonly
invented and commonly assigned patent applications, each of which
was filed on the same day or has the same priority date as the
present application, each of which is incorporated herein by
reference:
[0004] U.S. patent application Ser. No. 10/897,336 entitled
"Reducing Data Acquisition, Power and Processing for
Photoplethysmography and Other Applications", filed Jul. 21, 2004,
now allowed, and claiming the benefit of the earlier filing date of
U.S. patent application Ser. No. 10/895,004, filed Jul. 19,
2004;
[0005] U.S. patent application Ser. No. 10/895,165 entitled
"Reducing Data Acquisition, Power and Processing for Hemodynamic
Signal Amplitude Detection", filed Jul. 19, 2004, now
abandoned;
[0006] U.S. patent application Ser. No. 11/948,231 entitled
"Reducing Data Acquisition, Power and Processing for Hemodynamic
Signal Sampling", filed Nov. 30, 2007, now allowed;
[0007] U.S. patent application Ser. No. 10/895,004 entitled
"Reducing Data Acquisition, Power and Processing for Pulse Oximetry
Applications", filed Jul. 19, 2004, now abandoned; and
[0008] U.S. patent application Ser. No. 11/734,861 entitled
"Reducing Data Acquisition, Power and Processing for Pulse Oximetry
Applications", filed Apr. 13, 2007.
BACKGROUND OF THE INVENTION
[0009] FIG. 1 illustrates an exemplary photoplethysmography (PPG)
signal 102 produced using a PPG device (also known as a PPG
sensor). For timing reference, an electrocardiogram (ECG) signal
104 is also illustrated. The PPG signal 102 can be used to measure
the volume of arterial and venous vasculature. Additionally, a
measure of arterial pulse amplitude can be derived from the PPG
signal 102. A few tens to a few hundreds of milliseconds after the
QRS complex of the ECG signal 104, the PPG waveform reaches a
minimum and starts to increase. This is due to the increasing blood
volume in the arterioles as the systolic pulse reaches the
periphery. The delay is influenced by the distance that the PPG
sensor is placed from the heart. It requires approximately 100 msec
for the waveform to reach its maximum. The excursion from minimum
to maximum represents the arterial pulse amplitude. During
diastole, the recoil of the elastic arterial vessels continues to
force blood through the capillaries, so that blood flows through
the capillary bed throughout the entire cardiac cycle.
[0010] A PPG sensor (also called a pseudoplethysmography or
photoelectric plethysmography sensor) includes a light source and a
light detector. The PPG sensor utilizes the transmission or
reflection of light to demonstrate the changes in blood perfusion.
Such devices might be used, e.g., in the cardiology department or
intensive care department of a hospital or in a clinic for
diagnostic purposes related to vascular surgery.
[0011] A block diagram of an exemplary PPG sensor is shown in FIG.
2A. An exemplary mechanical arrangement for a noninvasive (i.e.,
not implanted) PPG sensor is shown in FIG. 2B. An exemplary
mechanical arrangement for a chronically implantable PPG sensor is
shown in FIG. 2C.
[0012] The PPG sensor includes a light source 206 and a light
detector 214. In one example, the light source 206 includes one or
more light-emitting diode (LED), although in alternative models an
incandescent lamp or laser diode can be used as the light source.
Referring to FIG. 2A, the light source 206 outputs a transmit light
signal 208 that is transmitted through and/or reflected by
(depending on the embodiment) patient tissue 210. For example,
light may be transmitted through a capillary bed such as in an
earlobe or finger tip. As arterial pulsations fill the capillary
bed and pre-capillary arterioles, the changes in volume of the
blood vessels modify the absorption, reflection and scattering of
the light. Stated another way, an arterial pulse in, for example, a
finger tip; or earlobe, causes blood volume to change, thereby
changing the optical density of the tissue. Therefore, the arterial
pulse modulates the intensity of the light passing through the
tissue.
[0013] A receive light signal 212 is received by the light detector
214. The light detector 214 can include, for example, a photodiode.
Changes in light intensity cause proportional changes in the
photodiode current, which can be converted to a varying analog
voltage light detection signal 216 by a transimpedance amplifier.
The light detector can, for example, alternatively include a
photoresistor, phototransistor, photodarlington or avalanche
photodiode. Light detectors are often also referred to as
photodetectors or photocells.
[0014] PPG sensors may operate in either a transmission
configuration or a reflection configuration. In the transmission
configuration, the light source 206 and the light detector 214 face
one another and a segment of the body (e.g., a finger or earlobe)
is interposed between the source 206 and the detector 214. In the
reflection configuration, the light source 206 and the light
detector 214 are mounted adjacent to one another, e.g., on the
surface of the body, as shown in FIG. 2B. In this configuration, a
fraction of light from the light source 206 is backscattered by the
tissue into the light detector 214.
[0015] Referring to FIG. 2C, if the PPG sensor is incorporated into
a chronically implantable device 220 (e.g., an implantable
cardioverter defibrillator (ICD), pacemaker, or any other
implantable device), the light source 206 and the light detector
214 can be mounted adjacent to one another on the housing or header
of the implantable device. The light source 206 and the light
detector 214 are preferably placed on the side of the implantable
device 220 that, following implantation, faces the chest wall, and
are configured such that light cannot pass directly from the source
to the detector. Thus, the reflection configuration is preferably
used when the plethysmography device is implemented in an
implantable device. The placement on the side of the device 220
that faces the chest wall maximizes the signal to noise ratio by 1)
directing the signal toward the highly vascularized musculature,
and 2) shielding the source and detector from ambient light that
enters the body through the skin. Alternatively, at the risk of
increasing susceptibility to ambient light, the light source 206
and the light detector 214 can be placed on the face of the device
that faces the skin of the patient. Additional details of an
implantable PPG device are disclosed in U.S. Pat. No. 6,491,639,
entitled "Extravascular Hemodynamic Sensor" (Turcott), which is
incorporated herein by reference.
[0016] The varying analog voltage light detection signal 216 that
is produced by the light detector 214 is a PPG signal. The PPG
signal is typically filtered, amplified and converted to a digital
signal using an analog to digital (A/D) converter (not necessarily
in the order). For example, the signal may be sampled at 500 Hz
(i.e., 500 samples per second) using a high resolution A/D
converter, and then the samples may undergo relatively intensive
post-acquisition filtering (e.g., using a 1000-point digital
filter). This relatively high sampling rate and relatively
intensive filtering consume battery power and processing resources.
While this may not be much of a concern with a non-implanted PPG
device (e.g., such as the one shown in FIG. 2B), minimizing power
consumption and processing is very important when it comes to
implantable devices. This is in part because invasive surgery is
required to replace the battery of an implanted device.
[0017] Accordingly, there is a desire to reduce, and hopefully
minimize, both the number of samples that are acquired, and the
associated processing of such samples. Additionally, there is a
desire to reduce the amount of power that is required to produce
and process the samples.
SUMMARY OF THE INVENTION
[0018] Embodiments of the present invention relate to methods, and
systems (which can be implemented as devices) for reducing the
amount of data, processing and/or power required to analyze
hemodynamic signals such as photoplethysmography signals and
arterial pressure signals.
[0019] In accordance with embodiments of the present invention, for
a window of time that spans at least two cycles of a cyclical body
function (e.g., heart beat or respiration), only one sample of a
hemodynamic signal is produced per cycle (e.g., cardiac cycle or
respiratory cycle), at a substantially same instant in each cycle.
This results in a plurality of samples being produced for the
window. The hemodynamic signal is then analyzed based on the
plurality of samples.
[0020] In accordance with embodiments of the present invention, for
each of a plurality of windows of time, only one sample of a
hemodynamic signal is produced per cycle of a cyclical body
function, at a substantially same instant in each cycle. This
results in a plurality of samples being produced for each window,
wherein each window spans at least two cycles of the cyclical body
function. The plurality of samples for each window can then be
averaged, to thereby produce an average value for each window. The
hemodynamic signal is then analyzed based on the average
values.
[0021] In accordance with further embodiments of the present
invention, in response to detecting a specific event associated
with a cyclical body function, analog circuitry is used to detect
and store a minimum and a maximum of the hemodynamic signal within
a window of time. For example, a first analog peak detector is used
to detect and store the maximum, and a second analog peak detector
is used to detect and store the minimum. The stored minimum and the
stored maximum are then sampled to produce a pair of samples from
which the peak-to-peak amplitude can be determined.
[0022] In accordance with other embodiments of the present
invention, in response to detecting a specific event associated
with a cyclical body function, a hemodynamic signal is continuously
sampled during a window following the detecting of the specific
event, wherein the window is shorter than a cycle associated with
the cyclical body function. The continuous sampling may be at about
20 Hz or greater if the cyclical body function is heart beat, or at
about 1 Hz or greater if the cyclical body function is respiration.
This results in a plurality of samples being produced for the
window. The hemodynamic signal is then analyzed based on the
plurality of samples.
[0023] Embodiments of the present invention also relate to reducing
the amount of processing required to determine blood oxygen (O2)
saturation levels.
[0024] In accordance with embodiments of the present invention, a
measure of DC offset and pulse amplitude associated with a received
first light signal (e.g., a red light signal) are obtained, and a
normalized first light pulse amplitude is produced therefrom.
Similarly, a measure of DC offset and pulse amplitude associated
with a received second light signal (e.g., an infrared or near
infrared light signal) is obtained, and a normalized second light
pulse amplitude is produced therefrom. Then, a two dimensional
look-up table is used to determine an O2 saturation level based on
the normalized first light pulse amplitude and the normalized
second light pulse amplitude.
[0025] In accordance with embodiments of the present invention,
light of a first wavelength and light of a second wavelength are
transmitted from a light source to a light detector (e.g., of a
pulse oximetry device), such that a corresponding DC offset and
pulse amplitude can be determined for light of the first wavelength
received at the light detector and a corresponding DC offset and
pulse amplitude can be determined for light of the second
wavelength received at the light detector. An intensity of the
transmitted light of the first wavelength is adjusted so that the
DC offset for the light of the first wavelength received at the
light detector is maintained at a substantially constant
predetermined level. Similarly, the intensity of the transmitted
light of the second wavelength is adjusted so that the DC offset
for the light of the second wavelength received at the light
detector is maintained at a substantially constant predetermined
level. This enables an O2 saturation level to be determined based
on a pulse amplitude determined for the light of the first
wavelength received at the light detector and a pulse amplitude
determined for the light of the second wavelength received at the
light detector, without having to normalize the pulse
amplitudes.
[0026] In accordance with an embodiment of the present invention,
prior to high pass filtering, a first light signal and a second
light signal are sampled to determine an estimate of each signal's
DC offset. The first light signal is indicative of light of a first
wavelength that is received at the light detector, and the second
light signal is indicative of light of a second wavelength that is
received at the light detector. Then, after high pass filtering,
the first light signal and the second light signal are sampled to
determine a pulse amplitude for each signal, wherein the sampling
before high pass filtering is at a lower frequency than the
sampling after high pass sampling, to thereby reduce the amount of
data produced. Then, an O2 saturation level is determined based on
the estimates of DC offset for the first and second light signals
and the pulse amplitudes for the first and second light
signals.
[0027] Other features and advantages of the invention will appear
from the following description in which the preferred embodiments
have been set forth in detail, in conjunction with the accompanying
drawings and claims.
BRIEF DESCRIPTION OF THE FIGURES
[0028] FIG. 1 illustrates exemplary PPG and ECG signals.
[0029] FIG. 2A is a high level block diagram of an exemplary PPG
sensor.
[0030] FIG. 2B is a simplified mechanical diagram illustrating a
portion of an exemplary PPG sensor.
[0031] FIG. 2C is a simplified mechanical diagram illustrating an
exemplary implantable PPG sensor.
[0032] FIGS. 3A and 3B illustrate exemplary light sources for use
in embodiments of the present invention.
[0033] FIG. 4 illustrates an exemplary light detector for use in
embodiments of the present invention.
[0034] FIG. 5 illustrates an overview of a monitoring system
according to an embodiment of the present invention.
[0035] FIG. 6 illustrates placement of an external telemetry unit
in, for example, a patient's bedroom.
[0036] FIG. 7 illustrates simultaneously recorded ECG and PPG
waveforms, which are useful for describing embodiments of the
present invention.
[0037] FIGS. 8A and 8B are high level flow diagrams useful for
describing embodiments of the present invention where only one
sample of a hemodynamic signal is produced per cycle of a cyclical
body function.
[0038] FIG. 9 is a high level flow diagram useful for describing
embodiments of the present invention where analog circuitry is used
to efficiently determine peak-to-peak amplitude of a hemodynamic
signal.
[0039] FIG. 10 is a high level flow diagram useful for describing
embodiments of the present invention where a hemodynamic signal is
continuously sampled during windows following specific events.
[0040] FIG. 11A illustrates an exemplary one dimensional look-up
table that could be used for determining oxygen saturation
levels.
[0041] FIG. 11B illustrates an exemplary two dimensional look-up
table, according to an embodiment of the present invention, that
could be used for determining oxygen saturation levels.
[0042] FIG. 12 is a high level flow diagram useful for describing
embodiments of the present invention where a two dimensional
look-up table, such as the one shown in FIG. 11B, is used to
determine oxygen saturation levels.
[0043] FIG. 13 is a high level flow diagram useful for describing
embodiments of the present invention in which lower frequency
sampling is used to determine estimates of DC offsets as compared
to the frequency of sampling used to determine measures of pulse
amplitude.
[0044] FIG. 14 is a high level flow diagram useful for describing
embodiments of the present invention where source optical power is
adjusted such that a predetermined DC level is detected at a light
detector, thereby eliminating the need to normalize measures of
pulse amplitude that are used to determine O2 saturation
levels.
[0045] FIG. 15A illustrates an exemplary implantable stimulation
device in electrical communication with a patient's heart by way of
three leads, which are suitable for delivering multi-chamber
stimulation and shock therapy.
[0046] FIG. 15B is a simplified block diagram of the multi-chamber
implantable stimulation device of FIG. 15A.
DETAILED DESCRIPTION OF THE INVENTION
Exemplary Photoplethysmography Sensors
[0047] As mentioned above, a PPG sensor includes a light source and
a light detector. FIGS. 3A and 3B illustrate exemplary light
sources for use in the embodiments of the present invention.
Referring first to FIG. 3A, exemplary light source 206 includes a
single LED that produces light signal 208. The LED can be, for
example, a model L53SRC/F red LED, or a model L53F3C infrared LED,
both manufactured by Kingbright Corporation, City of Industry,
California. Referring to FIG. 3B, a series of LEDs (e.g., LED1 and
LED2) can be used to increase the amount of optical power in light
signal 208. Separate LEDs can be used. Alternatively, dual emitter
combination LEDs can be used, such as model DLED-660/905-LL5-2,
manufactured by UDT Sensors, Inc., Hawthorne, Calif. In accordance
with an embodiment, a pair of separately driven LEDs are used,
where one of the LEDs is a red LED and the other is an infrared LED
(which can be a near infrared LED), collectively allowing for pulse
oximetry to be performed, providing for measures of blood oxygen
saturation. The light source 206 can be driven by one or more light
control signals 204, as shown in FIGS. 2A, 3A and 3B. In a
conventional PPG sensor, the transmit light signal 208 would have a
relatively constant average light intensity, though the light may
be pulsed rapidly. Accordingly, in a conventional PPG sensor, the
light control signal 204 is relatively constant when averaged over
a period of the pulse train.
[0048] One of ordinary skill in the art will appreciate that the
use of other LEDs and other light sources (e.g., a laser diode) are
within the spirit and scope of the present invention. Further, it
is possible that a green light (having a wavelength of about 530
nm) can be used instead of a red light.
[0049] Depending on the embodiment, the light source 206 may or may
not include additional elements that are used, for example, to
maintain a relatively constant current through an LED.
[0050] FIG. 4 illustrates an exemplary light detector for use in
embodiments of the present invention. Referring to FIG. 4, the
exemplary light detector 214 includes a photodiode PD operated in a
current sensing photoconductive mode feeding a transimpedance
amplifier 402. Photodiode PD can be, for example, a model
PIN-4.0-LLS, manufactured by UDT Sensors, Inc. The transimpedance
amplifier 402 includes a resistor R, a capacitor C and an
operational amplifier U, such as model ALD1701, manufactured by
Advanced Linear Devices, Inc., Sunnyvale, Calif. The amplifier 402,
including the RC circuit, performs low pass filtering and provides
gain. It also serves as an antialiasing filter if ND conversion is
applied directly to its output 216. One of ordinary skill in the
art will appreciate that a photodiode PD can alternatively be
operated in a voltage sensing configuration. Further, one of
ordinary skill in the art will appreciate that the use of other
photodiodes (e.g., an avalanche photodiode) and other light
detectors (e.g., a photoresistor, a photodarlington, a
phototransistor), are within the spirit and scope of the present
invention. One of ordinary skill in the art will also appreciate
that other amplifier configurations (e.g., an integrator amplifier
or a transistor based amplifier) can be used in place of the
transimpedance amplifier 402 shown in FIG. 4. An integrated
photodiode/amplifier (e.g., a Burr-Brown OPT101, available from
Burr-Brown Corporation, Tucson, Ariz.) can also be used.
[0051] In a conventional PPG sensor (e.g., FIG. 2B), a constant
average optical power is delivered by light source 206 (e.g., an
LED) and plethysmography information (e.g., measurements of the
waveform 102 shown in FIG. 1) is determined based on time varying
optical power incident on light detector 214. A PPG sensor device
can alternatively adjust the source of optical power such that a
relatively constant average light intensity is detected at a light
detector, as described in commonly assigned U.S. patent application
Ser. No. 09/907,349 (Turcott), filed Jul. 16, 2001, entitled
"Methods and Devices for Vascular Plethysmography Via Modulation of
Source Intensity," which is incorporated herein by reference. The
time-varying modulating signal (e.g., that controls the source
power) can then be used as the plethysmography signal the
information signal), rather than the time-varying detected optical
power. The time-varying detected optical power is used (e.g., in a
feedback loop) to adjust the source intensity.
[0052] FIG. 5 includes a block diagram that provides an overview of
a monitor 500, according to an embodiment of the present invention.
As will be explained in more detail below, the monitor 500 can be
used to analyze hemodynamic signals, such as PPG signals. The light
source 206 outputs a transmit light signal 208 of substantially
constant average light intensity (as controlled by light control
signal 204), though perhaps periodically or initially adjusted by
an automatic gain control feature so that the light detector 214 is
operating at a desirable point in its dynamic range. The light
signal 208 is transmitted through and/or reflected by (depending on
the embodiment) patient tissue 210. Receive light signal 212 is
received by the light detector 214. The light intensity of the
received light signal 212 is modulated by changes in blood volume
in patient tissue 210. The light detector 214 produces a light
detection signal 216 that is representative of the received light
signal 212. The light output signal 216, which is likely an analog
encoded information signal, is preferably filtered and amplified by
an analog signal processor block 522. A filtered and amplified
signal 524 is then provided to an analog to digital converted (A/D)
526, which provides a digital encoded plethysmography information
signal 528 to a microprocessor 530.
[0053] The microprocessor 530 analyzes the plethysmography signals
as represented by the encoded information signals 528. According to
embodiments of the present invention, the microprocessor 530
performs the averaging used in embodiments of the present
invention. The microprocessor 530 may also perform respiratory
monitoring, pacing interval optimization, etc., in accordance with
embodiments of the present invention.
[0054] If the monitor 500 is not implanted, the light source 206
and the light detector 214 can be made small and can conveniently
attach to a peripheral portion of the body, such as a finger, toe,
or ear. Thus, patients are likely to tolerate regular use of these
sensors for an extended period of time, such as during sleep each
night. Particular embodiments include a finger cuff, a wristband, a
configuration resembling a watch, and a configuration resembling a
clip-on earring. The light source 206 and light detector 214 could
be tethered to a larger unit containing the bulk of the electronic
circuitry (e.g., the microprocessor 530 and a memory 560). In this
case, the monitor 500 would be worn primarily when the patient is
sleeping. Alternatively, data (e.g., from the light detector 214,
ND 522, or microprocessor 530) could be continuously or
periodically be telemetered to a processor (e.g., the
microprocessor 530 or some other processor), which might be worn on
the patient's clothing or located in the patient's home and/or
office. In this case, the monitor could be worn both during sleep
and during activity. Nevertheless, despite the cost advantages of
an external embodiment, such an approach necessarily requires
patient cooperation. Because of the disadvantages associated with
this it may be preferable that the monitor 500 is an implanted
extravascular configuration. In addition, the monitoring function
just described can be integrated with a pacemaker or ICD in order
to enhance the therapy delivered by these devices. However, it
should be clear that many embodiments of the present invention are
not limited to implantable implementations.
[0055] The monitor 500 can also include a transmitter/receiver 550
(i.e., a telemetric circuit) and a memory 560. If the monitor 500
is chronically implanted, transmitter/receiver 550 enables the
operating parameters of the monitoring device 500 to be
non-invasively programmed into the memory 560 through telemetric
communications with an external device, such as a programmer or
transtelephonic transceiver. The transmitter/receiver 550, which is
preferably controlled by the microcontroller 530 (which is likely a
processor), also enables the monitor 500 to communicate with other
types of external processors. For example, the transmitter/receiver
550 enables plethysmography information and status information
relating to the operation of the device 500 (e.g., as contained in
the microcontroller 530 and/or memory 560) to be sent to an
external device (e.g., a remote processor or diagnostic system
analyzer) through an established communication link. The
microprocessor 530 can analyze a hemodynamic signal, and the
transmitter/receiver 550 can transmit the information to another
processor as appropriate. The transmitter/receiver 550 can
additionally, or alternatively, transmit waveform information to an
external device (e.g., a remote processor) that can analyze a
hemodynamic signal based on the information. Alternatively, the
encoded information signals (e.g., the light detection signal 216)
can be transmitted directly to an external device (e.g., a remote
processor), and the external device can perform appropriate
analysis.
[0056] For examples of a transmitter/receiver 550 (also known as a
telemetric circuit) of a chronically implantable device, see U.S.
Pat. No. 4,809,697, entitled "Interactive Programming and
Diagnostic System for use with Implantable Pacemaker" (Causey, III
et al.), and U.S. Pat. No. 4,944,299, entitled "High Speed Digital
Telemetry System for Implantable Device" (Silvian), each of which
is hereby incorporated herein by reference. Another example of a
telemetric circuit for use in a chronically implantable device is
the TR1000 transceiver manufactured by RF Monolithics, Dallas, Tex.
The TR 1000 is a single-chip, low-power, 916.5 MHz transceiver. The
operating frequency of 916.5 MHz is desirable because of the modest
requirements on antenna size it imposes.
[0057] The monitor 500 can also include a stimulation/alert block
540, that informs a patient, physician, clinician and/or any other
person (or processor) of the status of the patient. If the monitor
500 is implanted, an alert block 540 is preferably an external
device that telemetrically communicates with the microprocessor 530
(e.g., using transmitter/receiver 550). The stimulation/alert block
540 can include an indicator that provides, for example, an
acoustic, mechanical vibration, optical and/or electrical
indication and/or stimulation. Such an alert indicator can be
triggered when a criterion (e.g., threshold) is satisfied (e.g.,
exceeded), as discussed below. In one embodiment stimulation/alert
540 includes an inductive coil that generates both sound and
mechanical vibration. In an alternative embodiment, the function of
the stimulation/alert 540 is incorporated into the microprocessor
530 and the transmitter/receiver 550.
[0058] FIG. 6 illustrates placement of an external telemetry (i.e.,
transmitter/receiver) unit 602 in, for example, a patient's bedroom
or a physician's or clinician's office. The external telemetry unit
602, using telemetry at a distance, allows the transfer of data to
and from the monitor 500 if it is a chronically implanted device or
a device that clips on the finger, toe or earlobe, without the
active participation of the patient 604 or a clinician. The
external telemetry unit 602 is preferably positioned in a
location(s) regularly frequented by the patient, such as the
patient's bedroom, office, and/or automobile. The external
telemetry unit 602 can be in communication (e.g., through a
telephone line 606, network connection and/or wireless links) with
a central location for further processing or review (e.g., by a
clinician). Alternatively, the external telemetry unit can be in a
physician's or clinician's office so that data can be downloaded
from an implantable monitor 500 whenever the patient visits the
office. The data that is downloaded may have already been analyzed
by the implantable monitor, or the data that is downloaded can be
raw data that is analyzed after it is downloaded from the
implantable monitor.
[0059] A PPG sensor can use a single wavelength of light, or a
broad spectrum of many wavelengths. In the alternate embodiments,
the light source can be any source of radiant energy, including
laserdiode, heated filament, and ultrasound transducer. The
detector can be any detector of radiant energy, including
phototransistor, photodetector, ultrasound transducer,
piezoelectric material, and thermoelectric material.
Reducing Data Acquisition, Processing and/or Power Consumption
[0060] Embodiments of the present invention relate to reducing the
amount of data required (e.g., produced and/or stored) to analyze a
hemodynamic signal, such as a photoplethysmography (PPG) signal or
an arterial pressure signal. Embodiments of the present invention
also relate to reducing the amount of power consumption and
processing that is required to produce and/or analyze such data.
Embodiments of the present invention are further directed to
reducing the amount of data that may be stored for later analysis
of the data.
[0061] FIG. 7 illustrates simultaneously recorded ECG, and PPG
signals, labeled 702 and 712, respectively. In this plot, a
positive deflection of the PPG signal 712 is caused by increased
light absorption by the tissue, and a corresponding decrease in
detected light, as when a cardiac pulse causes an expansion of
peripheral vascular volume. Conversely, a negative deflection of
the PPG waveform results from a decrease in tissue light
absorption, and a corresponding increase in detected light, as when
vascular volume is reduced. Slow oscillation in the baseline of the
PPG signal 712 is due to positive-pressure ventilation, and likely
results from both the modulation of peripheral venous volume
induced by the changing intrathoracic pressure, and the modulation
of the arterial volume secondary to ventilation-induced changes in
arterial pressure, apparent in the waveform 712. Of lower amplitude
in this example, but still clearly apparent, are the pulsations in
the PPG waveform 712 due to the arrival of the cardiac pulse at the
periphery. Thus, effects arising from the modulation of both
arterial and venous vascular volumes can be seen in the raw PPG
signal 712.
[0062] Referring back to FIGS. 2, 4 and 5, the PPG signal 712 is an
example of the varying analog voltage light detection signal 216
that is produced by the light detector 214. As mentioned above,
such a PPG signal is typically filtered, amplified and converted to
a digital signal using an analog-to-digital (A/D) converter (not
necessarily in the order). For example, the signal may be sampled
at 500 Hz (i.e., 500 samples per second) using a high resolution
A/D converter, and then the samples may undergo relatively
intensive post-acquisition digital filtering (e.g., using a
1000-point filter). This relatively high sampling rate and
relatively intensive filtering consumes battery power and
processing resources. While this may not be much of a concern with
non-implanted PPG devices (e.g., such as the one shown in FIG. 2B),
minimizing power consumption and processing is very important when
it comes to implantable devices. This is in part because invasive
surgery is required to replace the battery of an implanted device.
Accordingly, there is a desire to reduce, and hopefully minimize,
both the number of samples that are acquired, and the associated
processing of such samples, which in turn will reduce and hopefully
minimize power consumption.
[0063] Producing One Sample Per Cycle of a Cyclical Body
Function
[0064] In accordance with embodiments of the present invention,
rather than continuously sampling at a high rate, the PPG signal
(e.g., signal 712) is sampled once for each heart beat. For
example, in accordance with an embodiment of the present invention,
the sampling of a PPG signal can be triggered by a sensed or paced
event. More specifically, the sensed event can be a ventricular or
atrial event, such as a contraction. Similarly, the paced event can
be a ventricular pace or an atrial pace. In some embodiments, the
PPG signal is sampled at a fixed delay after the sensed or paced
event.
[0065] For example, assume a patient's heart beat is 60 beats per
minute (i.e., 1 beat per second), and that sampling of the PPG
signal is 500 Hz (i.e., 500 samples per second). This would result
in 500 samples per heart beat. In contrast, with the just described
embodiment of the present invention, only one sample is obtained
per heart beat. Thus, the amount of acquired data is reduce by a
factor of 500.
[0066] Referring to FIG. 7, the open circles 714 represent samples
that were triggered by a ventricular contraction. In some
embodiments, the PPG signal 712 can be sampled at a fixed delay
after the ventricular event, as illustrated by the diamonds 716 in
FIG. 7.
[0067] The effects of cardiac pulsations are reduced by sampling
the signal at the same instant or point in each cardiac cycle
(which is preferably during diastole). Because the effects of
cardiac pulsations are sufficiently reduced, the need for filtering
of the sample is avoided (although filtering can still be performed
if desired). Further, the number of ND conversions is significantly
reduced (e.g., by a factor of 500 in this example). In addition, if
sample data is being stored for later analysis, the amount of
stored data is significantly reduced.
[0068] A systolic pulse does not reach a PPG sensor at a periphery
for approximately 200 milliseconds, resulting in a PPG signal being
essentially constant at the time of ventricular contraction. By
sampling the PPG signal during diastole, when the slope of the PPG
signal is small (rather than during the steep up slope during
systole), the effects of cardiac pulsations can be further
reduced.
[0069] The above described embodiments can be applied to
hemodynamic signals other than PPG signals. For example,
embodiments of the present invention can also be used to reduce the
amount of data processing and power required to analyze a pressure
signal. Such a pressure signal can be produced in various manners.
For example, a pressure catheter can be placed within an artery to
obtain an arterial pressure signal. Alternatively, a hollow lumen
catheter that is placed within an artery can be in communication
with an extravascular pressure transducer, thereby producing an
arterial pressure signal. In yet another alternative, a pressure
transducer can be placed on a pacing or defibrillation lead that is
positioned in the right ventricle, allowing an RV pressure signal
to be recorded. Such a pressure transducer can similarly be placed
in the right atrium, which would enable the acquisition of a right
atrial pressure signal. In still another alternative, thoracic
impedance or impedance of peripheral tissue can be used to assess
pulmonary or peripheral edema, respectively. In a further
alternative, thoracic impedance can be used to estimate cardiac
output and stroke volume, as is done in a commercially available
device produced by CardioDynamics, San Diego, Calif. These are just
a few examples, which are not meant to be limiting.
[0070] The above described embodiments of the present invention
will now be summarized and explained in further detail with
reference to the high level flow diagram of FIG. 8A. Referring to
FIG. 8A, at a step 802, for a window of time that spans at least
two cycles of a cyclical body function, one sample of the
hemodynamic signal is produced per cycle. In order to reduce (and
hopefully eliminate) the noise due to the cyclical body function,
the samples are produced at a substantially same instant in each
cycle. To further ensure that the effects of cardiac pulsations are
minimized, the samples can be produced during diastole, where the
slope of the hemodynamic signal is small, as has been explained
above.
[0071] Step 802 results in a plurality of samples for the window.
Next, at a step 804, the hemodynamic signal is analyzed based on
these plurality of samples. As mentioned above, the hemodynamic
signal can be, e.g., a PPG signal or an arterial pressure signal.
Additional details of step 804 are discussed below.
[0072] In accordance with embodiments of the present invention, the
cyclical body function referred to in step 802 is heart beat, and
the cycle referred to is a cardiac cycle. In other words, step 802
can be performed by producing one sample a hemodynamic signal per
cardiac cycle, at substantially the same instant in each cardiac
cycle, for a window of time that spans at least two cardiac
cycles.
[0073] In order to trigger the sampling at substantially the same
instant in each cardiac cycle, the sampling can be triggered in
response to a specific cardiac event, which can be detected based
on an ECG signal that is being simultaneously produced and
monitored. For example, in accordance with embodiments of the
present invention, the sampling is in response to sensing a
ventricular contraction. This can include sampling at the instant
the ventricular contraction is sensed, or a fixed delay after
sensing the ventricular contraction. In accordance with other
embodiments, the sampling is in response to a paced event, such as
ventricular pace, assuming the patient's heart is being paced. This
can include sampling at the instant of the ventricular pace, or a
fixed delay after the ventricular pace. Alternatively, sampling can
be in response to an atrial event, such as an atrial contraction or
an atrial pace. In a similar manner as described above, this may
include sampling when the atrial event is sensed/paced, or a fixed
delay thereafter.
[0074] Additional details of step 804 will now be described,
assuming the cyclical body function referred to in step 802 is
heart beat, and the cycle referred to is a cardiac cycle. In
accordance with embodiments of the present invention, step 804
includes monitoring respiration based on the plurality of samples
produced at step 802. For example, this can include determining a
rate of respiration based on the plurality of samples.
[0075] One way to accomplish this is to determine an average of the
plurality of samples, so that the average can serve as a threshold.
Then, the plurality of samples can be compared to the average to
thereby determine a number of threshold crossings. The rate of
respiration can then be determined based on the number of threshold
crossings, e.g., by counting the number of crossings from above the
threshold to below the threshold (or vice versa) for the window of
time, and converting that number to a conventional scale, such as
breaths per minute.
[0076] Another way to accomplish this is to determine an average of
the plurality of samples. Then, the plurality of samples can be
normalized by subtracting the average from each of the plurality of
samples, to thereby produce a plurality of normalized samples. The
plurality of normalized samples can then be compared to zero to
thereby determine a number of zero crossings. Then, in a similar
manner to that just described, the rate of respiration can be
determined based on the number of zero crossings.
[0077] In accordance with other embodiments of the present
invention, respiratory effort can be determined based on the
plurality of samples. This can be accomplished, e.g., by
determining a peak-to-peak amplitude based on the plurality of
samples. The peak-to-peak amplitude is indicative of the
respiratory effort in that an increase in peak-to-peak amplitude is
indicate of an increased respiratory effort, and a decrease in
peak-to-peak amplitude is indicative of a decrease in respiratory
effort.
[0078] Returning to the discussion of step 802, in alternative
embodiments the cyclical body function referred to in step 802 can
be respiration, and each cycle can be a respiratory cycle. In such
embodiments the sampling of the hemodynamic signal can be triggered
in response to sensing a respiratory event, such as the end or
beginning of inspiration or respiration. As with the previous
discussed embodiments, sampling can be triggered at the instant of
detecting a specific respiratory event, or a fixed delay after the
specific respiratory event. Such specific events may detected,
e.g., using measures of thoracic impedance which are often
determined by an implanted monitor and/or stimulation device. In
sampling once per respiratory cycle the variability in the PPG,
pressure, or other hemodynamic signal that is induced by
respiration or ventilation is avoided, and the need for filtering
is reduced as is the processing and volume of acquired data.
Eliminating respiratory variability is important, for example, in
obtaining relative estimates of average vascular volume from a PPG
signal, obtaining estimates of average pressure from a pressure
signal, or obtaining estimates of thoracic impedance (in order to
monitor for pulmonary edema) from a impedance measuring system.
These examples illustrate some of the applications of synchronously
sampling with a cyclical physiologic process, and are not meant to
be limiting.
[0079] As mentioned above, the hemodynamic signal that is being
sampled can be, e.g., a PPG signal, a pressure signal, or a signal
representative of pulmonary or peripheral edema, such as impedance.
Where the hemodynamic signal is a PRG signal, the PPG signal can be
produced using a light source and light detector that are not
implanted in a patient. However, embodiments of the present
invention are more likely to be implemented when the PPG signal is
produced using an implanted light source and light detector, which
are likely implemented as part of an implanted monitor and/or
stimulation device. This is because reducing amounts power,
processing and data storage is more important for implanted
devices.
[0080] A generally continuous PPG signal can be produced by a PPG
sensor, e.g., by continually driving the light source of the PPG
sensor, or pulsing the light source multiple times per cycle of the
cyclical body function. In accordance with embodiments of the
present invention just described above, the corresponding output of
the light detector of the PPG sensor is only sampled once per cycle
of the cyclical body function (e.g., once per cardiac cycle, or
once per respiratory cycle). Thus, the extent of sampling and the
number of samples produced (and possibly stored) is reduced. Also,
because the sampling is at substantially the point in each cycle,
there is no need to low pass filter the samples (for reasons
explained in detail above), further reducing processing and power
consumption.
[0081] Continually driving or pulsing the light source (i.e.,
driving the light source with voltage or current pulses) can
consume a significant amount of power. Accordingly, it may be
preferable to minimize the pulsing of the light source, and thus,
not produce a generally continuous PPG signal output. Rather, the
light source of the PPG sensor can be pulsed only once per cycle of
the cyclical body function (at substantially the same instant in
each cycle), causing the light detector to produce only one output
per cycle (at substantially the same instant in each cycle). This
is sufficient for the above discussed embodiments, since only one
sample per cycle need be produced, as has been described in detail
above. This will further reduce power consumption, because the
light source of the PPG sensor will consume less power if it is
driven less frequently. It is noted that step 802 is meant to
encompass this approach.
[0082] As mentioned above, rather than being a PPG signal, the
hemodynamic signal that is being analyzed can be an arterial
pressure signal. Such an arterial pressure signal can produced
using a pressure transducer placed within an artery. Alternatively,
the arterial pressure signal is produce using a hollow lumen
catheter that is placed within an artery and an extravascular
pressure transducer in communication with the catheter. In yet
another alternative, a pressure transducer is placed on a pacing or
defibrillation lead that is positioned in the right ventricle,
allowing an RV pressure signal to be recorded. In still another
alternative, thoracic impedance or impedance of peripheral tissue
can be used to assess pulmonary or peripheral edema,
respectively.
[0083] Depending on implementation, the time window referred to in
step 802 can span, e.g., a predetermined time interval, a
predetermined number of cardiac cycles, or a predetermined number
of respiratory cycles, any of which should be at least as long as
two cycles of the relevant cyclical body function.
[0084] At least two samples are produced at step 802, by producing
a sample only once per cycle. While such samples can be obtained
during consecutive cycles of the cyclical body function (e.g.,
during consecutive cardiac cycles), this is not required. That is,
during the sampling of the hemodynamic signal only once per cycle
at step 802, the cycles that are sampled need not be consecutive
cycles. For example, a sample could be produced every other cycle,
or even less uniformly and less frequently than that. In estimating
pulmonary edema, for example, one may elect to sample once a night.
While this may not be preferred, it is noted that step 802 is
intended to cover such approaches.
[0085] It is noted that step 802 can be repeated a number of times
(i.e., for a number of windows of time) before the samples produced
at step 802 are analyzed at step 804. For, example, sampling at
step 802 can be performed by an implanted device over a relatively
long period of time, and then the data (e.g., samples) can be
analyzed at a later time at step 804. The analysis at step 804 can
be performed by the implanted device, or the data obtained at step
802 can be downloaded (e.g., through telemetry) to an external
device that performs step 804. In this manner, with growing
interest in disease monitoring based on data obtained from
implantable devices (as well as external devices), embodiments of
the present invention can help alleviate problems associated with
acquiring large amounts of hemodynamic data over relatively long
periods of time between downloads to an external device.
[0086] The high level flow diagram of FIG. 8B will now be used to
provide additional details about embodiments where data is
collected for multiple windows of time.
[0087] Referring to FIG. 8B, at a step 812, for each of a plurality
of windows of time that each spans at least two cycles of the
cyclical body function, only one sample of a hemodynamic signal is
produced per cycle of a cyclical body function, at a substantially
same instant in each cycle. In this manner, a plurality of samples
is produced for each window. As was discussed above with reference
to step 802, during the sampling of the hemodynamic signal only
once per cycle at step 812, the cycles that are sampled within each
window need not be consecutive cycles. As with the discussion of
FIG. 8A the cycles could be, for example, cardiac cycles or
respiratory cycles. Similarly, the hemodynamic signal can be, for
example, a PPG signal, an arterial or right ventricular pressure
signal, or an impedance signal reflective of pulmonary or
peripheral edema. In a similar manner to that discussed above with
reference to step 802, the sampling can be in response to a
specific cardiac event or a specific respiratory event. This may
include sampling a fixed delay after the specific event.
[0088] The hemodynamic signal can then be analyzed based on the
samples produced at step 812. More specifically, at a step 814, for
each of the plurality of windows, the plurality of samples can be
averaged to thereby produce an average value for each window. At a
step 816, the hemodynamic signal can then be analyzed based on the
averages values. Other operations that characterize the amplitude
of the data can be used. For example, the data can be summed
without dividing by the number of data points that were included in
the sum. If this number is fixed then different sums can be
compared directly, and treated as if they were true averages.
Another alternative is to use the median of the plurality of
samples rather than the mathematical average. For simplicity we
refer to an average but this is not meant to imply that only a
precise mathematical average is acceptable. Rather, by `average` we
mean any characterization of the plurality of samples that
represents the ensemble properties.
[0089] In accordance with specific embodiments of the present
invention, the hemodynamic signal of which samples are produced in
step 812 is a PPG signal, and step 816 includes monitoring changes
in mean arterial pressure (MAP) based on changes in the average
values determined at step 814. Commonly invented and assigned U.S.
patent application Ser. No. 10/802,009, entitled "Methods, Systems
and Devices for Monitoring Mean Arterial Pressure" filed Mar. 15,
2004 (Attorney Docket No. A04P3003-US1), which is incorporated
herein by reference, explains how and why a PPG signal can be used
to monitor changes in mean arterial pressure. In the just mentioned
'009 application, each of a plurality of segments of a PPG signal
is averaged, to thereby produce a corresponding plurality of
average values, and changes in mean arterial pressure are monitored
based on changes in the average values. Embodiments of the present
invention can be used to improve upon such embodiments by reducing
the amount of data and processing used to monitor mean arterial
pressure.
[0090] In accordance with embodiments of the present invention,
monitoring the changes in mean arterial pressure includes
recognizing a change in the average values that corresponds to an
increase in arterial volume as an increase in mean arterial
pressure. Alternatively, or additionally, monitoring changes in
mean arterial pressure includes recognizing a change in the average
values that corresponds to a decrease in arterial volume as a
decrease in mean arterial pressure.
[0091] Further, pacing interval optimization can be based on the
monitored mean arterial pressure. For example, when comparing two
different AV delays, it can be concluded that the AV delay
producing the greatest increase in mean arterial pressure is the
better AV delay, meaning that it allows the heart to function with
greater mechanical efficiency. In one embodiment baseline pacing is
provided with a specific AV delay, and periodically the AV delay is
changed to test values with the pacing delivered at the test value
for a brief amount of time. The change in average arterial volume
is assessed by photoplethysmography as the changes in AV delay are
made. The AV delay that yields the greatest increase in blood
volume is then selected as the optimal AV delay. This is just one
example of how embodiments of the present invention can be used for
pacing optimization.
[0092] In accordance with other specific embodiments of the present
invention, the monitored changes in mean arterial pressure can be
used to select a type of anti-arrhythmia therapy, when an
arrhythmia is detected. Arrhythmias are irregular heartbeats that
feature either very rapid ventricular contractions (tachycardia),
an excessively slow heartbeat (bradychardia) or, commonly, extra or
"premature" beats. The most lethal arrhythmia is ventricular
fibrillation (VF), in which the ventricles undergo persistent and
disorganized activation. In this arrhythmia the heart is not
capable of pumping blood. Mean arterial blood pressure quickly
falls, and perfusion of the vital organs ceases. Once VF begins,
death will soon follow unless the arrhythmia is successfully
terminated. ICDs are generally quite effective at detecting VF
because of its exceedingly rapid electrical rate. Once VF is
detected, the ICDs are appropriately designed to deliver a
high-voltage shock, which is the most aggressive therapy. Some
arrhythmias do not necessarily require electrical therapy, e.g.,
atrial fibrillation (AF) when it is hemodynamically stable, that
is, when it doesn't compromise mean arterial pressure. Other
arrhythmias require electrical therapy, but it need not be
aggressive. For example, a low-rate ventricular tachycardia (VT) is
often hemodynamically stable. For these arrhythmias, low-voltage
anti-tachycardia pacing or low-energy cardioversion may terminate
the arrhythmia but consume less battery power and cause less
discomfort to the patient than the high-voltage shocks used to
terminate VF. Convention ICDs are typically programmed to deliver
aggressive therapy as quickly as possible for VF, to attempt to
terminate VT with less aggressive therapy, and to withhold
electrical therapy for AF. The problem is that for many rhythms,
the conventional ICD has no way of knowing whether the hemodynamic
status of the patient has been compromised because the same
electrical rhythm can have different hemodynamic consequences in
different patients, or in the same patient at different points in
time.
[0093] Thus, using embodiments of the present invention to detect
significant decreases in mean arterial pressure would allow an ICD
to select aggressive, high-voltage therapy only when necessary.
Similarly, recognizing that mean arterial pressure has not been
significantly compromised despite the detected onset of an
arrhythmia would allow the device to attempt less aggressive
techniques of arrhythmia termination. More specifically, at step
816, changes in average values can be compared to a threshold, and
a type of anti-arrhythmia therapy can be selected based on whether
the changes in the average values exceed the threshold. This may
include, for example, selecting a high voltage therapy if a change
in the average values exceeds the threshold, and selecting a lower
voltage therapy (or no electrical therapy) if a change in the
average values does not exceed the threshold.
[0094] Embodiments of the present invention can be used for pacing
interval optimization, as will now be discussed in further detail.
In accordance with embodiments of the present invention, the
hemodynamic signal referred to in step 812 is produced as a heart
is paced using a plurality of different sets of pacing intervals,
with each of the windows corresponds to a different one of the sets
of pacing intervals. Each set of pacing interval parameters can
include one or more pacing intervals (i.e., delays). The initiating
event, from which the interval/delay is specified, can be either a
delivered pace pulse, or a sensed depolarization. The pacing
interval parameters can be used, e.g., for multi-site pacing, and
may include, e.g., an atrio-ventricular (AV) delay, an
interventricular delay and/or an interatrial delay. Pacing
intervals can define an intra-chamber pacing delay or an
inter-chamber pacing delay. Pacing intervals can be used for two,
three or four chamber pacing. These are just a few examples, which
are not meant to limit the scope of the present invention.
[0095] By having each of the windows corresponds to a different one
of the sets of pacing intervals, a different average value (one for
each window) is produced for each set of pacing intervals, at step
814. Accordingly, step 816 can include performing pacing interval
optimization based on the average values, or more specifically,
based on changes in the average values. For example, step 816 can
include selecting one of the plurality of sets of pacing intervals,
as a preferred set, based on the average values. This may include,
e.g., selecting the set of pacing intervals that produced the
greatest average as the preferred set.
[0096] One of ordinary skill in the art will appreciate that at
steps 804 and 816 a hemodynamic signal could be analyzed in manners
other than those described in detail above, while still being
within the spirit and scope of the present invention.
[0097] Using Analog Circuitry to Detect Peak-to-Peak Amplitude
[0098] Embodiments of the present invention are also directed to
reducing the amount of sampling and processing required to detect
the peak-to-peak amplitude of a hemodynamic signal, such as a PPG
signal or an arterial pressure signal. Conventionally, such signals
would be continuously sampled to produce numerous samples from
which minimum and maximum amplitudes could be identified, and
peak-to-peak amplitude could then calculated based on the maximum
and minimum. Embodiments of the present invention, which will now
be described with reference to the high level flow diagram of FIG.
9, use analog circuitry to reduce the amount of sampling,
processing and power consumption required to detect peak-to-peak
amplitudes.
[0099] Referring to FIG. 9, at a step 902, in response to detecting
a specific event associated with a cyclical body function, analog
circuitry is used to detect and hold a minimum and a maximum of the
hemodynamic signal within a window of time. The analog circuitry
can include a first analog peak detector to detect and hold the
maximum, and a second analog peak detector to detect and hold the
minimum. Analog peak detectors are well known and thus need not be
described in further detail.
[0100] At a step 904, only the held minimum and the held maximum
are sampled by an A/D converter to produce a pair of digital
samples.
[0101] Then, at a step 906, a peak-to-peak amplitude is determined
based on the pair of samples.
[0102] As with previous embodiments discussed above, the cyclical
body function can be heart beat, and the specific event that
initiates the using of the analog circuitry can be, for example,
sensing a ventricular or atrial contraction, or a ventricular or
atrial pace. The cyclical body function can alternatively be
respiration, and the specific event that initiates the using of the
along circuitry can be expiration or inspiration. Similar to
previous embodiments, the use of the analog circuitry can be
triggered a fixed delay after a specific event.
[0103] Continuously Sampling During Short Windows
[0104] In accordance with the following embodiments of the present
invention, rather than simply producing only one sample per cycle
of a cyclical body function, the number of samples per cycle is
reduced by reducing the size of the window that is sampled. This
can be explained with reference back to FIG. 7 and with reference
to the high level flow diagram of FIG. 10. Conventionally, as
mentioned above, a hemodynamic signal that is to be analyzed would
be continuously sampled over multiple cycles of signal. Thus, by
only sampling the signal during short windows, the amount of
sampling (and number of samples produces) is reduced.
[0105] Referring to the flow diagram of FIG. 10, at a step 1002, in
response to detecting a specific event associated with a cyclical
body function, a hemodynamic signal (e.g., a PPG signal or an
arterial pressure signal) is continuously sampled during a window
following the detecting of a specific event, to thereby produce a
plurality of samples. In order to reduce the amount of samples, the
window should be shorter than a cycle associated with the cyclical
body function. For example, referring back to FIG. 7, the
horizontal lines 718 represent windows of time during which the PPG
signal 712 is continuously sampled. These windows are clearly
shorter than the cardiac cycles shown in the corresponding ECG
signal 702, yet they are long enough to include both the onset and
completion of the systolic cardiac pulse.
[0106] The windows should be wide enough so that the information of
interest is captured. For example, if the desire is to obtain the
systolic pulse amplitude of a PPG signal during each cardiac cycle,
then the window should span the portion of a cycle within which the
minimum and maximum are likely to be found. Within a cardiac cycle,
the minimum of a corresponding PPG signal will typically occur soon
after a ventricular contraction, followed soon after by a maximum
of the PPG signal. Accordingly, the window can span, for example, a
time beginning with a ventricular contraction and ending a fixed
time thereafter. Alternatively, the window can start a fixed delay
after a ventricular contraction, or some other specific event. In
still other embodiments, the temporal location of the window and
the length of the window can be dynamically adjusted to increase
the likelihood that the minimum and maximum are sampled, while also
minimizing the length of a window.
[0107] As with previous embodiments discussed above, the cyclical
body function can be heart beat, and the specific event that
initiates the sampling can be, for example, sensing a ventricular
or atrial contraction, or a ventricular or atrial pace. For those
embodiments where the cyclical body function is heart beat, the
continuous sampling is preferably performed at a rate of about 20
Hz or greater.
[0108] The cyclical body function can alternatively be respiration,
and the specific event that initiates the using of the analog
circuitry can be expiration or inspiration. Similar to previous
embodiments, the window in which sampling is continuously performed
can start a fixed delay after a specific event. For those
embodiments where the cyclical body function is respiration, the
continuous sampling is preferably performed at a rate of about 1 Hz
or greater.
[0109] Referring back to FIG. 10, at a next step 1004, the
hemodynamic signal is analyzed based on the plurality of samples
produced at step 1002. For example, step 1004 can include detecting
a peak-to-peak amplitude based on the plurality of samples.
[0110] In accordance with embodiments of the present invention,
step 1002 is repeated a plurality of times such that a plurality of
samples is produced for each of a plurality of windows. Then, step
1004 can include averaging the plurality of samples produced for
each of the windows to thereby produce an average value for each
window. In similar manners as were discussed in detail above, these
average values can be used, e.g., for monitoring mean arterial
pressure, for pacing interval optimization, and/or for selecting a
type of anti-arrhythmia therapy.
Measuring Blood Oxygen Saturation
[0111] Embodiments of the present invention are also directed to
reducing the amount of processing required to determine a blood
oxygen saturation (O2 saturation) level, which is more specifically
the percentage of hemoglobin that is saturated with oxygen.
Conventionally, measures of arterial O2 saturation are produced
using the well-known technique of pulse oximetry in the following
way: light of two different wavelengths, typically red (e.g., about
660 nm wavelength) and infrared or near infrared (e.g., about 940
nm wavelength), are alternately transmitted through or reflected by
patient tissue such that a single light detector receives incident
light that alternates between red and infrared light. More
specifically, one LED transmits red light and another LED transmits
infrared or near infrared light. The LEDs are serially pulsed to
produce an interleaved signal stream that is transmitted through or
reflected from tissue of a patient. As the light passes through
and/or is reflected from tissue, some of the energy is absorbed by
arterial and venous blood, tissue and the variable pulsations of
arterial blood. The interleaved red and infrared light stream is
received by the single light detector. The amplitudes of the red
light pulses in the light stream are differently effected by the
absorption than the infrared light pulses, with the absorptions
levels depending on the O2 saturation level of the blood.
[0112] Using electronic circuitry, firmware and/or software, the
received light signals in the infrared and red wavelengths are
analyzed so that O2 saturation levels can be determined. At a high
level, demultiplexing is used to produce a signal path for the
received red light and a separate signal path from the received
infrared light. Each signal path will typically include one or more
filters and an A/D converter to sample the received light signals.
The samples of the red light signal are then used to determine the
DC offset (i.e., average) and pulse amplitude of the received red
light. Similarly, the samples of the infrared light signal are then
used to determine the DC offset (i.e., average) and pulse amplitude
of the received infrared light. Each pulse amplitude is then
normalized (e.g., by dividing the pulse amplitude by the
corresponding DC offset) and a ratio of the red-to-infrared light
is determined by dividing the normalized red pulse amplitude by the
normalized infrared pulse amplitude. Then, a one dimensional
look-up table, such as the exemplary table of FIG. 11A, is
typically used to determine the O2 saturation level. A look-up
table is typically used because there is a well known one-to-one
correspondence between the red-to-infrared ratios and O2 saturation
levels. While the just described conventional scheme has worked
well, it would be beneficial if the amount of processing required
to obtain measures of O2 saturation levels could be reduced. It is
noted that for this embodiment and other embodiments described
herein it is possible that green light (having a wavelength of
about 530 nm) can be used instead of red light.
[0113] Using New Look-Up Tables to Determine O2 Saturation
[0114] According to embodiments of the present invention, rather
than calculating a ratio of red-to-infrared light (by dividing the
normalized red pulse amplitude by the normalized infrared pulse
amplitude), a two dimensional look-up table is used to determine an
O2 saturation level based on a normalized red pulse amplitude and a
normalized infrared pulse amplitude. Such a two dimensional look-up
table would have normalized red pulse amplitudes along a first axis
or dimension, and normalized infrared pulse amplitudes along a
second axis or dimension, with the cells of the table populated by
corresponding blood oxygenation levels. An example of such a
look-up table, according to an embodiment of the present invention,
is shown in FIG. 11B. By using a two dimensional look-up table, in
accordance with embodiments of the present invention, at least one
mathematical division operation is eliminated.
[0115] FIG. 12 is a high level flow diagram useful for describing
embodiments of the present invention where a two-dimensional
look-up table, such as the one shown in FIG. 11B, is used to
determine O2 saturation levels. Referring to FIG. 12, at a step
1202 a measure of DC offset and pulse amplitude associated with a
received first light signal is obtained, and a normalized first
light pulse amplitude is produced therefrom. The received first
light signal can be, e.g., a signal indicative of red light
received at a light detector of a pulse oximetry device.
[0116] At a step 1204, a measure of DC offset and pulse amplitude
associated with a received second light signal is obtained, and a
normalized second light pulse amplitude is produced therefrom. The
received second light signal can be, e.g., a signal indicative of
infrared or near infrared light received at the light detector of
the pulse oximetry device.
[0117] Then at a step 1206, a two dimensional look-up table (e.g.,
similar the table of FIG. 11B) is used to determine an O2
saturation level based on the normalized first light pulse
amplitude and the normalized second light pulse amplitude.
[0118] In accordance with another embodiment of the present
invention, a four dimensional look-up table (not shown) is used to
determine a blood oxygen saturation level based on non-normalized
red and infrared amplitudes and corresponding DC offsets. The four
axis or dimensions would include: non-normalized red pulse
amplitude, red DC offset, non-normalized infrared pulse amplitude
and infrared DC offset. By using a four dimensional look-up table,
in accordance with embodiments of the present invention, at least
three mathematical division operations are eliminated, as compared
to the conventional scheme described above. More specifically, the
four dimensional look-up table would be used to determine an O2
saturation level based on obtained measures of DC offset and pulse
amplitude associated with a received first light signal, and
measures of DC offset and pulse amplitude associated with a
received second light signal. The received first light signal can
be, e.g., a signal indicative of red light received at a light
detector of a pulse oximetry device. The received second light
signal can be, e.g., a signal indicative of infrared or near
infrared light received at the light detector of the pulse oximetry
device.
[0119] Simplifying Determinations of Dc Offset Used for Measuring
O2 Saturation
[0120] As was explained above, in order to determine measures of O2
saturation using pulse oximetry, measures of DC offset and pulse
amplitude should be obtained for the received red light and
infrared light. Typically, to determine the DC offset of each
signal, the signal is continuously sampled (prior to any high pass
filtering) to produce a plurality of samples from which an average
is determined, with the average being the DC offset. The pulse
amplitude of each signal is typically determined by high pass
filtering the signal to remove the DC offset, and then determining
the peak-to-peak amplitude of the signal (or the amplitude above
zero). Typically the DC offset and pulse amplitude are measured
using separate channels.
[0121] In contrast, in accordance with the following embodiments of
the present invention, a good estimate of DC offset is obtained
from as few as one sample of a DC coupled received light signal
(i.e., a signal that has not yet been high pass filtered to remove
DC components). More specifically, the magnitude of as few as one
sample of a DC coupled received light signal can be determined to
provide a good estimate of the DC offset. This is possible because
the magnitude of the DC offset is about 100 times larger than the
magnitude of the pulse amplitude variations, cardiac variations and
respiratory variations (e.g., the DC offset is measured in volts,
while the pulse amplitude is measured in millivolts). While as few
as one sample could be used to estimate DC offset, it is also
possible that a few samples could be averaged to provide a slightly
better estimate.
[0122] The measures of pulse amplitude can then be obtained in the
conventional manner, e.g., by high pass filtering a received light
signal and then continuously sampling the high passed filtered
signal to determine the amplitude. An improvement here is that the
DC offsets of received red and infrared signals are obtained with
less sampling, data acquisition, processing and power consumption.
In a broad sense, these embodiments can be characterized in that
the sampling prior to high pass filtering (for the purposes of
estimating DC offset) is at a lower frequency (and possibly a
significantly lower frequency) than the sampling frequency used to
measure pulse amplitude.
[0123] FIG. 13 will now be used to describe these embodiments of
the present invention. Referring to FIG. 13, at a step 1302, prior
to high pass filtering, a first light signal and a second light
signal are sampled to determine an estimate of each signal's DC
offset. The first light signal is indicative of light of a first
wavelength (e.g., red or green light) that is received at the light
detector. The second light signal is indicative of light of a
second wavelength (e.g., infrared or near infrared light) that is
received at the light detector. The first light signal can be
produced, e.g., by receiving red light at the light detector. The
second light signal can be produced, e.g., by receiving infrared or
near infrared light at the light detector.
[0124] At a step 1304, after high pass filtering, the first light
signal and the second light signal are sampled, at a higher
frequency than was used at step 1302, to determine a pulse
amplitude for each signal. For example, this can be accomplished by
continuous sampling the light signals multiple times per cardiac
cycle. Alternatively, analog peak detectors can be used, as was
described above, and only two samples could be produced per cardiac
cycle.
[0125] At a step 1306, an O2 saturation level is then determined
based on the estimates of DC offset for the first and second light
signal and the pulse amplitudes for the first and second light
signals. In accordance with an embodiment of the present invention,
the frequency of sampling at step 1302 is less than a corresponding
cardiac signal frequency, causing the light signals to be sampled
once or less per cardiac cycle. In accordance with an embodiment of
the present invention, the sampling at step 1302 is a frequency
that is at least half of the frequency of the sampling at step
1304, and can even be 100 times less than the frequency used at
step 1304.
[0126] Adjusting Source Optical Power
[0127] Embodiments of the present invention are also directed to
adjusting the source optical power such that the light detected at
the light detector of a pulse oximetry device has a substantially
stable predetermined DC offset. By doing this, the need to perform
normalization is eliminated, thereby reducing processing. Such
embodiments will be described with reference to the flow diagram of
FIG. 14.
[0128] Referring to FIG. 14, at a step 1402, light of a first
wavelength and light of a second wavelength are transmitted from a
light source to a light detector, such that a corresponding DC
offset and pulse amplitude can be determined for light of the first
wavelength received at the light detector and a corresponding DC
offset and pulse amplitude can be determined for light of the
second wavelength received at the light detector. The light of the
first wavelength can be, e.g., red or green light, and the light of
the second wavelength can be, e.g., infrared or near infrared
light.
[0129] The light source and light detector may be part of a pulse
oximetry device, which may or may not be implanted. If implanted,
the pulse oximetry device may be part of an implantable monitor or
stimulation device that performs other functions beside measuring
levels of O2 saturation. It is also possible that such a device
uses measures of O2 saturation for various purposes.
[0130] At a step 1404, an intensity of the transmitted light of the
first wavelength is adjusted so that the DC offset for the light of
the first wavelength received at the light detector is maintained
at a substantially constant predetermined level. Similarly, the
intensity of the transmitted light of the second wavelength is also
adjusted so that the DC offset for the light of the second
wavelength received at the light detector is maintained at a
substantially constant predetermined level. The predetermined level
for the light of the first wavelength may or may not be the same of
the predetermined level for the light of the second wavelength.
This step is most likely accomplished using feedback from the light
detector to the light source.
[0131] At a step 1406, an O2 saturation level is then determined
based on a pulse amplitude determined for the light of the first
wavelength received at the light detector and a pulse amplitude
determined for the light of the second wavelength received at the
light, detector. Because this step is performed without having to
normalize the pulse amplitudes, processing is reduced.
[0132] Exemplary Stimulation Device
[0133] Various embodiments discussed above relate to pacing
interval optimization, selecting a type of anti-arrhythmia therapy
in response to detecting an arrhythmia, sampling a signal in
response to detecting a cardiac or respiratory event, etc. For
completeness, an exemplary implanted stimulation device 1510 that
can be used to perform pacing, detect an arrhythmia, perform
anti-arrhythmia therapy, detect specific cardiac events, etc., is
described with reference to FIGS. 15A and 15B.
[0134] Referring to FIG. 15A, the exemplary implantable stimulation
device 1510 (also referred to as a pacing device, or a pacing
apparatus) is shown as being in electrical communication with a
patient's heart 1512 by way of three leads, 1520, 1524 and 1530,
suitable for delivering multi-chamber stimulation and shock
therapy. To sense atrial cardiac signals and to provide right
atrial chamber stimulation therapy, the stimulation device 1510 is
coupled to an implantable right atrial lead 1520 having at least an
atrial tip electrode 1522, which typically is implanted in the
patient's right atrial appendage. Stimulation device 1510 can be
integrated with one of the embodiments of the monitor 500 discussed
above. That is, a common housing can be used to contain the
elements of the monitor 500 (e.g., a light source 206 and light
detector 214) and the elements of the stimulation device 1510. More
generally, the sensor that produces an arterial plethysmography
signal can be in the same housing that contains the stimulation
device. Alternatively, separates housings can be used to house the
monitor 500 and the stimulation device 1510. This is of course
necessary if a monitor in not implantable (e.g., in embodiments
where the one or more sensors associated with a monitor are
incorporated into a finger cuff, a wristband, a configuration
resembling a watch, or a configuration resembling a clip-on
earring). As mentioned above, the sensor that produces the arterial
plethysmography signal can be a PPG sensor. Alternatively, a strain
gauge, a linear displacement sensor, or ultrasound transducer can
be used.
[0135] Referring to FIG. 15A, to sense left atrial and ventricular
cardiac signals and to provide left-chamber pacing therapy, the
stimulation device 1510 is coupled to a "coronary sinus" lead 1524
designed for placement in the "coronary sinus region" via the
coronary sinus for positioning a distal electrode adjacent to the
left ventricle and/or additional electrode(s) adjacent to the left
atrium. As used herein, the phrase "coronary sinus region" refers
to the vasculature of the left ventricle, including any portion of
the coronary sinus, great cardiac vein, left marginal vein, left
posterior ventricular vein, middle cardiac vein, and/or small
cardiac vein or any other cardiac vein accessible by the coronary
sinus.
[0136] A pressure transducer can be located on the right atrial
lead 1520, on lead 1530, or on an a separate lead (now shown), to
enable the device 1510 to produce a right atrial, right
ventricular, or arterial pulse pressure hemodynamic signal,
respectively. Alternatively, pressure transducers can be placed
from right heart leads transeptally into the left atrium or even
the left ventricle in order obtain left atrial and left ventricular
pressures. It is also possible that a hollow lumen catheter can be
inserted in an artery or within a heart chamber, with the hollow
lumen catheter being in communication with a pressure transducer
located within the device housing 1540. These approaches to
pressure sensing can be used in a chronically implanted device, or
can be placed temporarily to allow acute measurements, as during
diagnostic or therapeutic maneuvers or for monitoring in intensive
care settings.
[0137] The exemplary coronary sinus lead 1524 is designed to
receive atrial and ventricular cardiac signals and to deliver left
ventricular pacing therapy using at least a left ventricular tip
electrode 1526, left atrial pacing therapy using at least a left
atrial ring electrode 1527, and shocking therapy using at least a
left atrial coil electrode 1528.
[0138] The stimulation device 1510 is also shown in electrical
communication with the patient's heart 1512 by way of an
implantable right ventricular lead 1530 having, in this embodiment,
a right ventricular tip electrode 1532, a right ventricular ring
electrode 1534, a right ventricular (RV) coil electrode 1536, and
an SVC coil electrode 1538. Typically, the right ventricular lead
1530 is transvenously inserted into the heart 1512 so as to place
the right ventricular tip electrode 1532 in the right ventricular
apex so that the RV coil electrode 1536 will be positioned in the
right ventricle and the SVC coil electrode 1538 will be positioned
in the superior vena cava. Accordingly, the right ventricular lead
1530 is capable of receiving cardiac signals and delivering
stimulation in the form of pacing and shock therapy to the right
ventricle.
[0139] As illustrated in FIG. 15B, a simplified block diagram is
shown of the multi-chamber implantable stimulation device 1510,
which is capable of treating both fast and slow arrhythmias with
stimulation therapy, including cardioversion, defibrillation, and
pacing stimulation. While a particular multi-chamber device is
shown, this is for illustration purposes only, and one of skill in
the art could readily duplicate, eliminate or disable the
appropriate circuitry in any desired combination to provide a
device capable of treating the appropriate chamber(s) with
cardioversion, defibrillation and pacing stimulation.
[0140] The housing 1540 for the stimulation device 1510, shown
schematically in FIG. 15B, is often referred to as the "can",
"case" or "case electrode" and may be programmably selected to act
as the return electrode for all "unipolar" modes. The housing 1540
may further be used as a return electrode alone or in combination
with one or more of the coil electrodes, 1528, 1536 and 1538, for
shocking purposes. The housing 1540 further includes a connector
(not shown) having a plurality of terminals, 1542, 1544, 1546,
1548, 1552, 1554, 1556, and 1558 (shown schematically and, for
convenience, the names of the electrodes to which they are
connected are shown next to the terminals). As such, to achieve
right atrial sensing and pacing, the connector includes at least a
right atrial tip terminal (A.sub.R TIP) 1542 adapted for connection
to the atrial tip electrode 1522.
[0141] To achieve left chamber sensing, pacing and shocking, the
connector includes at least a left ventricular tip terminal
(V.sub.L TIP) 1544, a left atrial ring terminal (A.sub.L RING)
1546, and a left atrial shocking terminal (A.sub.L COIL) 1548,
which are adapted for connection to the left ventricular tip
electrode 1526, the left atrial ring electrode 1527, and the left
atrial coil electrode 1528, respectively.
[0142] To support right chamber sensing, pacing and shocking, the
connector further includes a right ventricular tip terminal
(V.sub.R TIP) 1552, a right ventricular ring terminal (V.sub.R
RING) 1554, a right ventricular shocking terminal (R.sub.V COIL)
1556, and an SVC shocking terminal (SVC COIL) 1558, which are
adapted for connection to the right ventricular tip electrode 1532,
right ventricular ring electrode 1534, the RV coil electrode 1536,
and the SVC coil electrode 1538, respectively. At the core of the
stimulation device 1510 is a programmable microcontroller 1560
which controls the various modes of stimulation therapy, including
pacing optimization and anti-arrhythmia therapy. As is well known
in the art, the microcontroller 1560 typically includes a
microprocessor, or equivalent control circuitry, designed
specifically for controlling the delivery of stimulation therapy
and can further include RAM or ROM memory, logic and timing
circuitry, state machine circuitry, and I/O circuitry. Typically,
the microcontroller 1560 includes the ability to analyze signals
(data) as controlled by a program code stored in a designated block
of memory. The details of the design of the microcontroller 1560
are not critical to the present invention. Rather, any suitable
microcontroller 1560 can be used to carry out the functions
described herein. The use of microprocessor-based control circuits
for performing timing and data analysis functions are well known in
the art. In specific embodiment of the present invention, the
microcontroller 1560 performs some or all of the steps associated
with detecting specific events, triggering sampling, monitoring
mean arterial pressure, pacing interval optimization and selecting
an appropriate anti-arrhythmia therapy. It is noted that the
microcontroller 1560 and microprocessor 530 can be one in the same,
or separate, depending on implementation and embodiment.
[0143] Representative types of control circuitry that may be used
with the invention include the microprocessor-based control system
of U.S. Pat. No. 4,940,052 (Mann et. al.) and the state-machines of
U.S. Pat. No. 4,712,555 (Sholder) and U.S. Pat. No. 4,944,298
(Sholder). For a more detailed description of the various timing
intervals used within the stimulation device and their
inter-relationship, see U.S. Pat. No. 4,788,980 (Mann et. al.). The
'052, '555, '298 and '980 patents are incorporated herein by
reference.
[0144] As shown in FIG. 15B, an atrial pulse generator 1570 and a
ventricular pulse generator 1572 generate pacing stimulation pulses
for delivery by the right atrial lead 1520, the right ventricular
lead 1530, and/or the coronary sinus lead 1524 via an electrode
configuration switch 1574. It is understood that in order to
provide stimulation therapy in each of the four chambers of the
heart, the atrial and ventricular pulse generators, 1570 and 1572,
may include dedicated, independent pulse generators, multiplexed
pulse generators, or shared pulse generators. The pulse generators,
1570 and 1572, are controlled by the microcontroller 1560 via
appropriate control signals, 1576 and 1578, respectively, to
trigger or inhibit the stimulation pulses.
[0145] The microcontroller 1560 further includes timing control
circuitry 1579 which is used to control pacing parameters (e.g.,
the timing of stimulation pulses) as well as to keep track of the
timing of refractory periods, PVARP intervals, noise detection
windows, evoked response windows, alert intervals, marker channel
timing, etc., which is well known in the art. Examples of pacing
parameters include, but are not limited to, atrio-ventricular
delay, interventricular delay and interatrial delay.
[0146] The switch bank 1574 includes a plurality of switches for
connecting the desired electrodes to the appropriate I/O circuits,
thereby providing complete electrode programmability. Accordingly,
the switch 1574, in response to a control signal 1580 from the
microcontroller 1560, determines the polarity of the stimulation
pulses (e.g., unipolar, bipolar, combipolar, etc.) by selectively
closing the appropriate combination of switches (not shown) as is
known in the art.
[0147] Atrial sensing circuits 1582 and, ventricular sensing
circuits 1584 may also be selectively coupled to the right atrial
lead 1520, coronary sinus lead 1524, and the right ventricular lead
1530, through the switch 1574 for detecting the presence of cardiac
activity in each of the four chambers of the heart. Accordingly,
the atrial (ATR. SENSE) and ventricular (VTR. SENSE) sensing
circuits, 1582 and 1584, may include dedicated sense amplifiers,
multiplexed amplifiers, or shared amplifiers. The switch 1574
determines the "sensing polarity" of the cardiac signal by
selectively closing the appropriate switches, as is also known in
the art. In this way, the clinician may program the sensing
polarity independent of the stimulation polarity.
[0148] Each sensing circuit, 1582 and 1584, preferably employs one
or more low power, precision amplifiers with programmable gain
and/or automatic gain control, bandpass filtering, and a threshold
detection circuit, as known in the art, to selectively sense the
cardiac signal of interest. The automatic gain control enables the
device 1510 to deal effectively with the difficult problem of
sensing the low amplitude signal characteristics of atrial or
ventricular fibrillation. Such sensing circuits, 1582 and 1584, can
be used to determine cardiac performance values used in the present
invention.
[0149] The outputs of the atrial and ventricular sensing circuits,
1582 and 1584, are connected to the microcontroller 1560 which, in
turn, are able to trigger or inhibit the atrial and ventricular
pulse generators, 1570 and 1572, respectively, in a demand fashion
in response to the absence or presence of cardiac activity, in the
appropriate chambers of the heart. The sensing circuits, 1582 and
1584, in turn, receive control signals over signal lines, 1586 and
1588, from the microcontroller 1560 for purposes of measuring
cardiac performance at appropriate times, and for controlling the
gain, threshold, polarization charge removal circuitry (not shown),
and timing of any blocking circuitry (not shown) coupled to the
inputs of the sensing circuits, 1582 and 1586.
[0150] For arrhythmia detection, the device 1510 utilizes the
atrial and ventricular sensing circuits, 1582 and 1584, to sense
cardiac signals to determine whether a rhythm is physiologic or
pathologic. The timing intervals between sensed events (e.g.,
P-waves, R-waves, and depolarization signals associated with
fibrillation which are sometimes referred to as "F-waves" or
"Fib-waves") are then classified by the microcontroller 1560 by
comparing them to a predefined rate zone limit (i.e., bradycardia,
normal, low rate VT, high rate VT, and fibrillation rate zones) and
various other characteristics (e.g., sudden onset, stability,
physiologic sensors, and morphology, etc.) in order to assist with
determining the type of remedial therapy that is needed (e.g.,
bradycardia pacing, anti-tachycardia pacing, cardioversion shocks
or defibrillation shocks, collectively referred to as "tiered
therapy").
[0151] Cardiac signals are also applied to the inputs of an
analog-to-digital (ND) data acquisition system 1590. The data
acquisition system 1590 is configured to acquire intracardiac
electrogram signals, convert the raw analog data into a digital
signal, and store the digital signals for later processing and/or
telemetric transmission to an external device 1502. The data
acquisition system 1590 is coupled to the right atrial lead 1520,
the coronary sinus lead 1524, and the right ventricular lead 1530
through the switch 1574 to sample cardiac signals across any pair
of desired electrodes.
[0152] Advantageously, the data acquisition system 1590 can be
coupled to the microcontroller 1560, or other detection circuitry,
for detecting an evoked response from the heart 1512 in response to
an applied stimulus, thereby aiding in the detection of "capture".
Capture occurs when an electrical stimulus applied to the heart is
of sufficient energy to depolarize the cardiac tissue, thereby
causing the heart muscle to contract. The microcontroller 1560
detects a depolarization signal during a window following a
stimulation pulse, the presence of which indicates that capture has
occurred. The microcontroller 1560 enables capture detection by
triggering the ventricular pulse generator 1572 to generate a
stimulation pulse, starting a capture detection window using the
timing control circuitry 1579 within the microcontroller 1560, and
enabling the data acquisition system 1590 via control signal 1592
to sample the cardiac signal that falls in the capture detection
window and, based on the amplitude, determines if capture has
occurred.
[0153] The implementation of capture detection circuitry and
algorithms are well known. See for example, U.S. Pat. No. 4,729,376
(Decote, Jr.); U.S. Pat. No. 4,708,142 (Decote, Jr.); U.S. Pat. No.
4,686,988 (Sholder); U.S. Pat. No. 4,969,467 (Callaghan et. al.);
and U.S. Pat. No. 5,350,410 (Mann et. al.), which patents are
hereby incorporated herein by reference. The type of capture
detection system used is not critical to the present invention.
[0154] The microcontroller 1560 is further coupled to a memory 1594
by a suitable data/address bus 1596, wherein the programmable
operating parameters used by the microcontroller 1560 are stored
and modified, as required, in order to customize the operation of
the stimulation device 1510 to suit the needs of a particular
patient. Such operating parameters define, for example, pacing
pulse amplitude, pulse duration, electrode polarity, rate,
sensitivity, automatic features, arrhythmia detection criteria, and
the amplitude, waveshape and vector of each shocking pulse to be
delivered to the patient's heart 1512 within each respective tier
of therapy.
[0155] A feature of the present invention is the ability to sense
and store data (e.g., from the data acquisition system 1590). Such
data can then be used for subsequent analysis to guide the
programming of the device and/or to monitor mean arterial pressure,
appropriately adjust pacing interval parameters, select optimum
pacing intervals, and/or select appropriate anti-arrhythmia
therapy, in accordance with embodiments of the present
invention.
[0156] Advantageously, the operating parameters of the implantable
device 1510 may be non-invasively programmed into the memory 1594
through a telemetry circuit 1501 in telemetric communication with
the external device 1502, such as a programmer, transtelephonic
transceiver, or a diagnostic system analyzer. The telemetry circuit
1501 is activated by the microcontroller by a control signal 1506.
The telemetry circuit 1501 advantageously allows intracardiac
electrograms and status information relating to the operation of
the device 1510 (as contained in the microcontroller 1560 or memory
1594) to be sent to an external device 1502 through an established
communication link 1504.
[0157] For examples of such devices, see U.S. Pat. No. 4,809,697,
entitled "Interactive Programming and Diagnostic System for use
with Implantable Pacemaker" (Causey, III et al.); U.S. Pat. No.
4,944,299, entitled "High Speed Digital Telemetry System for
Implantable Device" (Silvian); and U.S. patent application Ser. No.
09/223,422, filed Dec. 30, 1998, entitled "Efficient Generation of
Sensing Signals in an Implantable Medical Device such as a
Pacemaker or ICD" (note: this relates to transfer of EGM data)
(McClure et al.), which patents are hereby incorporated herein by
reference.
[0158] In accordance with an embodiment, the stimulation device
1510 further includes one or more physiologic sensors 1508, that
can be used, e.g., to produce an arterial plethysmography signal.
The physiologic sensors 1508 can include, for example, a PPG sensor
having light source and a light detector (e.g., similar to light
source 206 and light detector 214). In other words, portions of the
PPG sensor, described in detail above, can be incorporated into or
with the stimulation device 1510. This would enable the stimulation
device 1510 to produce an arterial plethysmography signal that is
useful for monitoring the mean arterial pressure of the patient and
obtaining data for pulse oximetry calculations, and to produce a
venous plethysmography signal that is useful for monitoring
respiration. The microcontroller 1560 can respond by selecting
and/or adjusting the various pacing parameters (e.g.,
atrio-ventricular delay, interventricular delay, interatrial delay
etc.), based on measures of mean arterial pressure determined using
the plethysmography signal. As explained above, other types of
sensors can alternatively be used to produce an arterial
plethysmography signal.
[0159] The microcontroller 1560 can respond to changes in mean
arterial pressure by adjusting the various pacing parameters in
accordance with the embodiments of the present invention. The
microcontroller 1560 controls adjustments of pacing parameters by,
for example, controlling the stimulation pulses generated by the
atrial and ventricular pulse generators, 1570 and 1572. While shown
as being included within the stimulation device 1510, it is to be
understood that the physiologic sensor 1508 may also be external to
the stimulation device 1510, yet still be implanted within or
carried by the patient. More specifically, the sensor 1508 can be
located inside the device 1510, on the surface of the device 1510,
in a header of the device 1510, or on a lead (which can be placed
inside or outside the bloodstream).
[0160] The stimulation device 1510 additionally includes a battery
1511 which provides operating power to all of the circuits shown in
FIG. 15B. For the stimulation device 1510, which employs shocking
therapy, the battery 1511 must be capable of operating at low
current drains for long periods of time, and then be capable of
providing high-current pulses (for capacitor charging) when the
patient requires a shock pulse. The battery 1511 must also have a
predictable discharge characteristic so that elective replacement
time can be detected. Accordingly, the device 1510 preferably
employs lithium/silver vanadium oxide batteries, as is true for
most (if not all) current devices.
[0161] The stimulation device 1510 further includes a magnet
detection circuitry (not shown), coupled to the microcontroller
1560. It is the purpose of the magnet detection circuitry to detect
when a magnet is placed over the stimulation device 1510, which
magnet may be used by a clinician to perform various test functions
of the stimulation device 10 and/or to signal the microcontroller
1560 that the external programmer 1502 is in place to receive or
transmit data to the microcontroller 1560 through the telemetry
circuits 100.
[0162] As further shown in FIG. 15B, the device 1510 is shown as
having an impedance measuring circuit 1513 which is enabled by the
microcontroller 1560 via a control signal 1514. The known uses for
an impedance measuring circuit 1513 include, but are not limited
to, lead impedance surveillance during the acute and chronic phases
for proper lead positioning or dislodgement; detecting operable
electrodes and automatically switching to an operable pair if
dislodgement occurs; measuring respiration or minute ventilation;
measuring thoracic impedance for determining shock thresholds;
measuring thoracic impedance for detecting and assessing the
severity of pulmonary edema; detecting when the device has been
implanted; measuring stroke volume; and detecting the opening of
heart valves, etc. The impedance measuring circuit 1513 is
advantageously coupled to the switch 1574 so that any desired
electrode may be used. In addition, to facilitate the measurement
of peripheral tissue edema, extra electrodes can be added to the
device housing, thereby limiting the test electric field to the
peripheral tissue.
[0163] In the case where the stimulation device 1510 is also
intended to operate as an implantable cardioverter/defibrillator
(ICD) device, it must detect the occurrence of an arrhythmia, and
automatically apply an appropriate electrical shock therapy to the
heart aimed at terminating the detected arrhythmia. To this end,
the microcontroller 1560 further controls a shocking circuit 1516
by way of a control signal 1518. The shocking circuit 1516
generates shocking pulses of low (up to 0.5 Joules), moderate
(0.5-10 Joules), or high energy (11 to 40 Joules), as controlled by
the microcontroller 1560. Such shocking pulses are applied to the
patient's heart 1512 through at least two shocking electrodes, and
as shown in this embodiment, selected from the left atrial coil
electrode 1528, the RV coil electrode 1536, and/or the SVC coil
electrode 1538. As noted above, the housing 1540 may act as an
active electrode in combination with the RV electrode 1536, or as
part of a split electrical vector using the SVC coil electrode 1538
or the left atrial coil electrode 1528 (i.e., using the RV
electrode as a common electrode).
[0164] Cardioversion shocks are generally considered to be of low
to moderate energy level (so as to minimize pain felt by the
patient), and/or synchronized with an R-wave and/or pertaining to
the treatment of tachycardia. Defibrillation shocks are generally
of moderate to high energy level (i.e., corresponding to thresholds
in the range of 5-40 Joules), delivered asynchronously (since
R-waves may be too disorganized to be recognized), and pertaining
exclusively to the treatment of fibrillation. Accordingly, the
microcontroller 1560 is capable of controlling the synchronous or
asynchronous delivery of the shocking pulses. Another approach to
electrical anti-arrhythmia therapy is anti-tachycardia pacing, in
which low-voltage pacing pulses are applied to pace-terminate the
arrhythmia. This approach is particularly effective in low rate
ventricular tachycardias.
CONCLUSION
[0165] The previous description of the preferred embodiments is
provided to enable any person skilled in the art to make or use the
embodiments of the present invention. While the invention has been
particularly shown and described with reference to preferred
embodiments thereof, it will be understood by those skilled in the
art that various changes in form and details may be made therein
without departing from the spirit and scope of the invention.
* * * * *